journal of the mechanical behavior of biomedical materials 40 (2014) 127–139

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Research Paper

Mechanical properties and in vivo performance of load-bearing fiber-reinforced composite intramedullary nails with improved torsional strength N. Moritza,b, N. Strandberga, D.S. Zhaoa, R. Mattilab, L. Paracchinic, P.K. Vallittub, H.T. Aroa,n a Orthopaedic Research Unit, Department of Orthopedic Surgery and Traumatology, University of Turku and University Central Hospital, FI-20520 Turku, Finland b Turku Clinical Biomaterials Centre - TCBC, Department of Biomaterials Science, Institute of Dentistry, University of Turku, FI-20520 Turku, Finland c INGEO Snc., Varallo Pombia, NO, Italy

art i cle i nfo

ab st rac t

Article history:

Fiber-reinforced composites (FRC) could be feasible materials for fracture fixation devices if

Received 25 May 2014

the mechanical properties of the composites are congruent with the local structural

Received in revised form

properties of bone. In a recently developed FRC implant, bisphenol A dimethacrylate

16 August 2014

(BisGMA) and triethylene glycol dimethacrylate (TEGDMA) resin was reinforced with

Accepted 22 August 2014

unidirectional E-glass fibers. The addition of a braided glass fiber sleeving to the

Available online 1 September 2014

unidirectional fibers increased the torsional strength (99.5 MPa) of the FRC implants at

Keywords:

the expense of the flexural strength (602.0 MPa). The flexural modulus was 15.3 GPa. Two

Fiber-reinforced composites

types of FRC intramedullary nails were prepared; first type was FRC as such, second type

Orthopedic surgery

was FRC with a surface layer of bioactive glass (BG) granules. Experimental oblong

Implant

subtrochanteric defect was created in 14 rabbits. The defect, which reduced the torsional

Mechanical testing

strength of the bones by 66%, was fixed with an FRC intramedullary nail of either type. The

Intramedullary nail

contralateral intact femur served as the control. This model simulated surgical stabilization of bone metastasis. After 12 weeks of follow-up, the femurs were harvested and analyzed by torsional testing, micro-CT and hard tissue histology. Healed undisplaced periimplant fractures were noticed in half of the animals irrespective of the type of FRC implant. Torsional testing showed no significant differences between the implantation groups. The torsional strength of the bones stabilized by either type of FRC implant was 83% of that of the contralateral femurs. In histological analysis, no implant debris and no

n Correspondence to: Orthopaedic Surgery and Traumatology, Turku University Hospital, T-hospital, Room E115, FI-20521, Turku, Finland. Tel.: þ358 40 353 7644; fax: þ358 2 313 8111. E-mail address: hannu.aro@utu.fi (H.T. Aro).

http://dx.doi.org/10.1016/j.jmbbm.2014.08.020 1751-6161/& 2014 Elsevier Ltd. All rights reserved.

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journal of the mechanical behavior of biomedical materials 40 (2014) 127 –139

adverse tissue reactions were observed. While the mechanical properties of the modified FRCs were suboptimal, the FRC intramedullary nails supported the femurs without structural failure, even in the cases of peri-implant fractures. & 2014 Elsevier Ltd. All rights reserved.

1.

Introduction

Metallic fracture fixation devices and joint replacement implants have mechanical properties different from those of bone. The mismatch of the mechanical properties leads to bone stress-shielding with adverse remodeling processes. While the structural stiffness of a metallic implant could be altered e.g. by the reduction of the cross-section of the device, the capability of an implant to withstand the dynamic physiological loading remains the key issue. In this sense, fiber-reinforced composites (FRC) are attractive alternative materials for load-bearing implants (Evans and Gregson, 1998); FRCs possess good fatigue resistance while the mechanical properties of the orthopedic FRC implants can be tailored to closely match with the local structural properties of the implantation site. In addition, non-metallic FRC implants may produce clinically fewer artefacts with CT and MRI imaging, and allow the use of postoperative radiation therapy without risk of scattering of radiation. The typical shortcomings of FRCs include the decrease of strength due to water sorption and the leaching of residual monomers or hydrolysis products into the peri-implant space. Early research on FRCs intended for orthopedic implants was focused on epoxy matric reinforced with carbon fibers (Hastings, 1978; Tayton et al., 1982; Ali et al., 1990). Other polymer matrices were also investigated (Ramakrishna et al., 2001). The choice of carbon fibers for reinforcement was motivated by a higher stiffness/mass ratio of carbon fibers compared to that of glass fibers (Hastings, 1978). However, attempts to introduce implants made of carbon FRCs into clinical practice were hampered by the reports of inadequate performance of these implants (Allcock and Ali, 1997; Adam et al., 2002) and susceptibility to generate carbon fiber debris (Gillett et al., 1985; Kettunen et al., 1999; Jockisch et al., 1992; Minovic et al., 2001). Successful use of thermoset polymer matrices reinforced with E-glass fibers in dental applications (Vallittu, 1999; Vallittu and Sevelius, 2000) prompted a renewed interest in FRCs as potential candidate materials for skeletal reconstructions. Clinical use of implants made of this type of FRC material has already started in cranial reconstructions (Aitasalo et al., 2014). The new concept of non-metallic FRC implants amenable for use of biomimetic and antibacterial surfaces led to the large-scale EU-project (NewBone, www.hb. se/ih/polymer/newbone). As a crude biomechanical and biological testing of the FRC concept, we first evaluated a simple FRC intramedullary nail made of unidirectional glass fibers (Zhao et al., 2009) and now we report the mechanical properties and in vivo behavior of novel FRC intramedullary nails with the composite structure made of unidirectional fiber core and braided biaxial glass fiber sleeving.

2.

Materials and methods

2.1.

Research methodology of the study

In our previous study, FRC intramedullary nails were surgically implanted in the rabbit femurs (Zhao et al., 2009). The implants were reinforced with continuous unidirectional glass fibers. Under physiological loading conditions, the FRC intramedullary nails were able to support the bones without structural damage to the implants. Performance of the FRC implants was comparable to that of the control implants made of titanium. However, two animals showed undisplaced healed peri-implant fractures in the femurs with FRC implants, which were probably due to the low torsional strength of the FRC intramedullary nails. It was assumed that the addition of a braided biaxial glass fiber sleeving (Hiermer et al., 1998) to the reinforcement component of the FRC could improve the torsional strength of the intramedullary nails. To test this assumption, two types of composite specimens were compared. Specimens coded UF, were reinforced with continuous unidirectional fibers. These specimens were similar to the FRC type used in our previous study (Zhao et al., 2009). Specimens of the second material, coded UFS, were reinforced with continuous unidirectional E-glass fibers in conjunction with a braided biaxial glass fiber sleeving (Fig. 1). The mechanical properties of the specimens were assessed in three-point bending and torsion tests. Finite element (FE) modeling of the FRC specimens was also performed and compared with the experimental data of the mechanical testing. For the in vivo study, two types of FRC intramedullary nails were prepared in the geometry reported earlier (Zhao et al., 2009). The implants were coded UFS and UFS-BG. UFS intramedullary nails were made of UFS material. UFS-BG intramedullary nails were made of UFS material with bioactive glass (BG) granules attached to the surface of the FRC. BG was expected to improve the osteointegration of the implants. Previously reported rabbit model was employed (Zhao et al., 2009). The selected animal model simulated surgical stabilization of bone metastasis in the subtrochanteric femur region by means of the FRC intramedullary nails. Torsional strength of cadaver bones stabilized with UFS implants was assessed prior to the implantation in the animals.

2.2.

Preparation of FRC specimens

Rod-shaped (∅¼ 3.3 mm) 70 mm long specimens were prepared for the mechanical tests. The photopolymerisable resin consisted of bisphenol A dimethacrylate (BisGMA) and triethylene glycol dimethacrylate (TEGDMA) co-polymers with the BisGMA/TEGDMA ratio of 70/30 wt% with the

journal of the mechanical behavior of biomedical materials 40 (2014) 127 –139

129

Fig. 1 – Conceptual illustration of FRC implants with BG as the surface component: (a) UF-type reinforced with long unidirectional fibers (Zhao et al. 2009) and (b) UFS-BG reinforced with long unidirectional E-glass fibers in conjunction with a biaxial glass fiber sleeving. UF and UFS specimens as well as UFS intramedullary nails did not have the BG granules. (c) MicroCT-based reconstruction of the actual biaxial glass fiber sleeving used in the UFS specimens and implants; the braid angle was 271 in the ready composite.

Table 1 – Summary of materials used for preparation of FRC specimens and implants. Material

Type of material

Manufacturer

Bisphenol A dimethacrylate (BisGMA)

Co-monomer

Triethylene glycol dimethacrylate (TEGDMA)

Co-monomer

2-(Dimethylamino)ethyl methacrylate (DMAEMA)

Activator

Camphorquinone

Photoinitiator

Tex 2400, Hybons 2002

Reinforcment component: E-glass fiber roving Reinforcment component: braided biaxial E-glass sleeving Osteoconductive surface component

Röhm Chemische Fabrik GmbH, Darmstadt, Germany Aldrich Chemie GmbH, Steinheim, Germany Fluka Chemie GmbH, Buchs, Switzerland Sigma-Aldrich GmbH, Buchs, Switzerland PPG Industries, Pittsburgh, USA A&P Technologies Inc., Cincinnati, USA Vivoxid Ltd., Turku, Finland

Silasox™ E26L25X Bioactive glass (BG) S53P4, granules: 315–500 mm fraction, oxide composition (wt%): SiO2 53%, Na2O 23%, CaO 20%, P2O5 4%

addition of camphorquinone (0.7 wt%) and 2-(dimethylamino)ethyl methacrylate (DMAEMA) (0.7 wt%). In UF specimens, the BisGMA/TEGDMA copolymer matrix was reinforced with unidirectional pre-impregnated E-glass fibers (Fig. 1a). The fibers were aligned along the long axis of the specimens. In UFS specimens, a braided biaxial glass fiber sleeving was wrapped around the unidirectional fibers (Fig. 1b). The unidirectional fibers and the braided biaxial sleeving were used as obtained from the manufacturers. No surface modification, such as silane-treatment, was performed. To keep the same diameter in both groups of specimens, the amount of unidirectional fibers was less in UFS specimens. The specimens were first pre-cured with blue light for 40 s with a hand-held unit (Optilux 501, Kerr, Danbury, USA). Subsequently, the specimens were cured in a vacuum light oven (Visio Beta

vario, 3M/ESPE, Seefeld, Germany) for 15 min and in a light oven (Liculite, Dentsply De Trey GmbH, Dreieich, Germany) for 90 min at an ambient temperature. Raw materials used for the preparation of the specimens are listed in Table 1.

2.3.

Mechanical testing of FRC specimens

Three-point bending test of the FRC specimens was performed according to standard ASTM D790-98. A universal material testing machine (Lloyd Instruments LRX, Lloyd Instruments Ltd., Fareham, UK), was employed in the testing. Six specimens of UF and UFS were tested. The loading nose and supports had cylindrical shape; the span between the supports was 50 mm. The speed of the loading nose was kept 1 mm/min. Details of the experimental setup are given in

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journal of the mechanical behavior of biomedical materials 40 (2014) 127 –139

Fig. 2 – Schematic illustration of the specimens prepared for (a) three-point bending and (b) torsion tests. In the threepoint bending test the diameter of the bearings was 6 mm. Fig. 2a. The flexural stress (σf) and the flexural modulus (EB) were calculated using following equations: σf ¼

EB ¼

8Fl

ð1Þ

πd3 4ml3 3πd4

ð2Þ

where F is the maximum load, l is the span length, d is the diameter of the specimen, m is the slope of the tangent to the initial linear portion of the load–deflection curve. Testing of the FRC specimens in torsion was performed according to standard SFS-ISO 458-1 using a universal material testing machine (Avalon Technologies, Rochester, MI, USA). Six specimens of UF and UFS were tested. To prevent the damage to the specimens due to the attachment to the jaws of the testing equipment, both ends of the specimens were press-fit into 17.5 mm long pieces of brass tubing. Thus the gauge length of the specimen was 35 mm. The torsional loading was applied at the speed of 221/min. Details of the experimental setup are given in Fig. 2b. The material properties related to torsion, i.e. shear stress (τ) and shear modulus (G) were calculated as follows: τ¼

2MT πr3

ð3Þ

G¼

2lMT απr4

ð4Þ

where MT is the torque, r is the radius of the specimen and α is the angle of rotation. Torsional stiffness (S) was calculated by the linear fit at the linear section of the torque-angle curve.

2.4.

Finite element modeling

The purpose of the FE modeling was to find the parameters (Young's modulus and Poisson ratio) which would give the

best match between the experimental data obtained in the actual mechanical testing of the specimens and the FE simulations of these experiments. These parameters could be used in the future FE modeling of different FRC implants. Computer-aided design (CAD) software (SolidWorks 2007 SP 4.0 Dassault Systèmes, U.S.) was used to create the three-dimensional CAD models of the UF and UFS specimens. These models were made using the dimensions of the actual specimens. The internal structure of the CAD models mimicked the structure of the actual specimens; long unidirectional fiber core in the case of UF and long unidirectional fiber core in conjunction with the biaxial sleeving in the case UFS. The thickness of the sleeving and the orientation of the fibers were measured from the micro-CT images of the specimens (Fig. 1c). NEiFusion 1.2 and NEiNastran 9.1 software (Noran Engineering Inc., U.S.) were used in the FE analysis. The CAD models were further processed to generate the FE meshes. Node tetrahedral solid parabolic formulation and node hexahedral solid parabolic formulation were used in the case of UF meshes. In the case of UFS meshes, node layered structural solid were used in addition to node tetrahedral solid parabolic and node hexahedral solid parabolic formulations. The UF mesh contained 38,221 elements and the UFS contained 127,199 elements. In the first phase of the FE analysis, all materials were considered linear, homogenous and isotropic. Later, the increment of the mathematical models was necessary to describe non-linear, homogenous and orthotropic materials. The FE simulations mimicked the actual three-point bending and torsional tests performed in the laboratory. Comparisons of the results of the FE simulations with the experimental data obtained in the laboratory tests served as the benchmarks in the FE modeling process.

2.5.

Preparation of FRC intramedullary nails

Twenty FRC intramedullary nails were prepared. Implants had cylindrical shape (90 mm, ∅ ¼3.3 mm). To fit to the medullary cavity of the rabbit femur, the implants were slightly curved: the deflection of the middle of the implant was 4 mm, as shown in Fig. 3a. The photopolymerisable resin consisted of BisGMA/TEGDMA (70/30 wt%) with camphorquinone (0.7 wt%) and 2-dimethylamino ethylmethacrylate (0.7 wt%). The resin was reinforced with unidirectional preimpregnated E-glass fibers in conjunction with a braided biaxial glass fiber sleeving (Fig. 1b). An open Teflon mold was used to attain the desired geometry of the FRC implants. The implants were first pre-cured with blue light for 40 s with a hand-held unit (Optilux 501, Kerr, Danbury, USA). The UFS intramedullary nails were prepared by this method. For the UFS-BG implants, an additional thin layer of resin was applied to the surface of the specimens. Thereafter, the implants were rolled over the BG granules spread on a flat surface. BG granules were used as obtained from the manufacturer. No surface modification, such as silane-treatment, was performed. The application of surface layer of BG granules was followed by pre-curing for 40 s with the handheld unit. Subsequently, the implants were cured in a vacuum light oven (Visio Beta vario, 3M/ESPE, Seefeld, Germany) for 15 min and in a light oven (Liculite, Dentsply De Trey GmbH, Dreieich, Germany) for 90 min at an ambient

journal of the mechanical behavior of biomedical materials 40 (2014) 127 –139

131

Fig. 3 – (a) Shape and dimensions of the implants as determined by bone geometry; (b) position and shape of the experimental bone defect. (c) Sites of the hard-tissue sections.

temperature. Raw materials used for the preparation of the specimens are listed in Table 1. Implants were sterilized in the autoclave at 121 1C for 20 min (pressure 0.1 MPa) immediately prior to surgery. Six of the UFS-type FRC intramedullary nails were used for the mechanical tests with cadaver femurs. Fourteen UFS intramedullary nails (n ¼7) and UFS-BG intramedullary nails (n¼ 7) were used in the animal study.

2.6. Mechanical testing of cadaver femurs with implants in torsion Cadaver rabbit femurs were harvested from intra-institutional donors euthanized for other reasons. The bones were cleaned off the soft tissues, wrapped in saline-moist tissues and closed in plastic bags before immediate freezing at 20 1C. The bones were thawed, re-wrapped in moist tissues and kept on ice in a closed polystyrene container until tested. Once thawed, the specimens were tested without further refreezing. The experimental set-up for the mechanical testing of cadaver femurs with implants was described earlier by Zhao and colleagues (Zhao et al., 2009). The experimental setup simulated the post-operative stage of the animal experiment. In brief, an oblong defect was created in the anterior aspect of the femur shaft (Fig. 3b). The defect was stabilized by the insertion of the FRC intramedullary nail (Fig. 3c). Intact contralateral femur of the same animal served as control. Six pairs of bones were tested in torsion. The proximal and distal ends of the femurs were embedded in polymethylmetacrylate (PMMA) blocks and attached to the testing machine with pointed screws. Lloyd Instruments LRX universal material testing machine was used as a rotary actuator attached to a custom-made fixture. The load on left and right femurs was applied in opposite directions, to simulate the normal

physiological loading of the femurs. Care was taken to keep the bones moist during the preparation and testing. Loading in torsion was performed at the rate of 641/min.

2.7.

Animal surgery

The animal study protocol was approved by the Provincial State Office of Western Finland (permit #2007-00763). All experiments were carried out in accordance with the national guidelines for animal welfare. Fourteen adult male New Zealand white rabbits (Harlan, Horst, the Netherlands) with an average weight of 4 kg (range 3.3–4.8 kg) were used. Atropine (1 mg/kg subcutaneously) was given as a premedication and the anesthesia was induced by fentanyl–fluanisone combination (0.3 mL/kg intramuscularly, Hypnorms, Janssen Pharmaceuticals, UK). The animals received a single dose of benzylpenicillin as a standard prophylactic antibiotic (500,000 IU intramuscularly). Under standard sterile surgical conditions, the right proximal femur was exposed via an anteromedial intermuscular approach and the oblong defect (3  9 mm2) was created in the anterior aspect of the femur shaft (Fig. 3b). Subsequently, the medullary cavity was opened in the trochanteric fossa, the canal was reamed and the FRC intramedullary nail was introduced. At the end of the surgery, naloxone (0.1 mg/kg intramuscularly, Narcantis, Du Pont Pharmaceuticals, UK) was given to reverse the anesthesia. After surgery, functional activity of the animals was not restricted. For three days, the animals received post-operative pain medication of caprofen (1.5 mg/kg subcutaneously, Rimadyls vet. Pfizer Oy, Espoo, Finland). After 12 weeks, the animals were euthanized with an intravenous administration of sodium pentobarbital (Mebunat, Orion Oyj, Espoo, Finland). The femurs were harvested and analyzed

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journal of the mechanical behavior of biomedical materials 40 (2014) 127 –139

using biomechanical testing in torsion, micro-CT imaging, and hard tissue histology.

2.8.

Mechanical testing of the retrieved femurs in torsion

The operated femurs stabilized with UFS and UFS-BG implants and contralateral intact femurs were tested in torsion as fresh specimens within 4 h after retrieval. The bones were carefully cleaned of remaining soft tissues and the ends of the femurs were embedded in PMMA. The femurs were wrapped in moist tissues and kept on ice in a closed polystyrene container until tested. The proximal and distal ends of the femurs embedded in PMMA blocks were attached to the testing machine with pointed screws. Lloyd Instruments LRX universal material testing machine was used as an axial actuator attached to a custom-made fixture. The load was applied to the left and right femurs in opposite directions, to simulate the normal physiological loading of the femurs. Care was taken to keep the bones moist during the preparation and testing. Loading in torsion was performed at the rate of 641/min. After the testing, the bone specimens were placed in 40% ethanol for further processing.

2.9.

Micro-CT and hard tissue histology

The retrieved femurs stabilized with UFS and UFS-BG implants were imaged by a micro-CT scanner (Skyscan 1072, Skyscan N.V. Kontich, Belgium). To fit the imaging space of the scanner, the bones were cut in three sections. After the imaging, the bone specimens were prepared for histological analysis. The specimens were dehydrated in a graded series of ethanol, cleared in xylene, and embedded in isobornylmethacrylate (Technovit 1200 VLC, Kulzer, Germany). Sections of 20 mm thickness were prepared with a cutting and grinding technique (Exakt Apparatebau, Hamburg, Germany) and stained by the van Gieson method for light microscopy. Four histological sections were prepared from each femur as shown in Fig. 3c.

2.10.

Statistical analysis

Normal distribution of the data was analyzed using Kolmogorov–Smirnov test. Unpaired Student's t-test was used to compare the mechanical properties of UF and UFS specimens. Paired Student's t-test was used to compare the maximum shear stress of cadaver femurs stabilized with the UFS

intramedullary nails with those of the intact contralateral femurs. In the analysis of the mechanical data of the in vivo experiment, paired Student's t-test was used for the intraanimal comparison of the stabilized and contralateral intact femurs and unpaired Student's t-test for the inter-animal comparison of the femurs stabilized with a UFS or UFS-BG implant. The level of statistical significance was considered to be 0.05. Statistical analysis was performed using IBM SPSS Statistics program (version 19, SPSS Inc., USA).

3.

Results

3.1.

Mechanical testing of FRC specimens

The mechanical properties of UF and UFS specimens were assessed in three-point bending and torsion tests. In both types of tests, there were two distinct areas in the respective load–displacement and torque–angle curves. In the beginning, there was a linear increase in the load or torque to a certain level, when the surface cracks started to appear and the slope of the curve changed. These flexural stress and shear stress values, denoted σf1 and τ1 respectively, were considered of primary importance for the performance of the FRC implants. With further increase in the load or torque, the ultimate failure of the specimens occurred. These maximum flexural stress (flexural strength) and maximum shear stress (torsional strength) values were denoted σf2 and τ2 respectively. For UFS specimens, in both types of tests, the surface cracks appeared at an earlier stage of testing than in the case of UF specimens. The measured flexural properties of the implant materials are presented in Table 2, the torsional properties of the implant materials are presented in Table 3. In the three-point bending test, the flexural stress values (σf1 and σf2) of the UFS specimens were significantly decreased compared with those of UF specimens (Table 2). However, the modulus of elasticity in bending, EB, remained at the same level of around 16 GPa, close to the modulus of elasticity of bone. As expected,due to the addition of a bidirectional fiber sleeving to the reinforcement component, the shear stress values of the UFS specimens were significantly increased compared with those of UF specimens (Table 3). In addition, the UFS specimens started to exhibit surface cracks at the rotational angle of 291 and ultimately failed at the rotational angle of 1601, while the UF specimens started to crack at the rotational angle of 531 and could be rotated to almost 3601

Table 2 – Flexural properties of FRC specimens (sample mean7sample standard deviation). Parameter

Flexural stress at the start of surface cracking, σf1, [MPa] Maximum flexural stress (flexural strength), σf2, [MPa] Flexural modulus, EB [GPa] n

Unpaired Student's t-test. ns ¼not significant.

nn

Statistical analysisn

FRC specimens UF n¼8

UFS n ¼6

745.57131.9 802.07112.7 17.073.3

278.2747.6 602.0781.6 15.371.8

po0.001 p¼ 0.003 nsnn

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journal of the mechanical behavior of biomedical materials 40 (2014) 127 –139

Table 3 – Torsional properties of FRC specimens (sample mean7sample standard deviation). Parameter

Torque at the start of surface cracking, MT1 [Nm] Shear stress at the start of surface cracking, τ1 [MPa] Angle of rotation at the start of surface cracking, α1 [deg] Maximum torque, MT2 [Nm] Maximum shear stress (torsional strength), τ2 [MPa] Angle of rotation at ultimate failure, α2 [deg] Shear modulus, G [GPa] Torsional stiffness, S [Nm/deg] n

Statistical analysisn

FRC specimens UF n¼6

UFS n¼6

0.370.1 25.573.8 52.9711.4 0.4970.06 39.275.0 353.879.5 0.570.1 0.0170.01

0.770.1 53.5710.5 29.474.5 1.2570.18 99.5714.7 160.1730.3 1.970.6 0.0370.01

po0.001 po0.001 p ¼0.002 po0.001 po0.001 po0.001 po0.001 p ¼0.004

Unpaired Student's t-test.

Table 4 – Biomechanical testing of the retrieved femurs with FRC implants, intra-animal comparison (sample mean7 sample standard deviation). Parameter

Implantation group UFS n ¼7

UFS-BG n¼7

Maximum torque [Nm] Healed femur with FRC intramedullary nail Contralateral intact femur Statistical analysisn

4.170.4 5.070.4 po0.001

4.370.4 5.270.3 p ¼0.002

Torsional stiffness [Nm/deg] Healed femur with FRC intramedullary nail Contralateral intact femur Statistical analysisn

0.2670.07 0.2270.04 p¼ 0.027

0.3070.06 0.2470.02 p ¼0.014

Angle of rotation at ultimate failure [deg] Healed femur with FRC intramedullary nail Contralateral intact femur Statistical analysisn

17.873.3 24.675.1 p¼ 0.003

16.772.9 22.571.3 p ¼0.002

n

Paired Student's t-test.

Table 5 – Biomechanical testing results of the retrieved femurs with FRC implants expressed as the percentage of the contralateral intact femur value (sample mean7sample standard deviation). Parameter

Maximum torque, ratio [%] Torsional stiffness, ratio [%] Angle of rotation at ultimate failure, ratio [deg] n

Statistical analysisn

Implantation group UFS n¼7

UFS-BG n ¼7

82.776.1 73.0711.4 119.7711.5

83.077.91 74.4712.9 126.7721.0

nsnn ns ns

Unpaired Student's t-test. ns ¼ not significant.

nn

with the delamination of the structure but without the failure of the fibers.

TEGDMA resin, Young's modulus (E) was 3.7 GPa and Poisson ratio (ν) was 0.37 while for E-glass fibers, Young's modulus (E) was 72 GPa and Poisson ratio (ν) was 0.32.

3.2.

Finite element modeling

The following material properties provided the best fit between FE analysis and laboratory tests: for BisGMA/

There was a reasonable match of the selected material models with the data obtained from the actual mechanical tests of UF and UFS specimens. The FE analysis confirmed the differences in the mechanical properties of UF and

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journal of the mechanical behavior of biomedical materials 40 (2014) 127 –139

Fig. 4 – Micro-CT-based three-dimensional reconstruction of a femur with the FRC-BG intramedullary nail: (a) Cortical defect 12 weeks after surgery. Spiral fracture line due to biomechanical testing in torsion is seen passing the lower corner of the defect at a site of high stress-concentrations (Zhao et al., 2009). (b) Three-dimensional reconstruction showing endosteal bone formation around the intramedullary implant.

Fig. 5 – Micro-CT-based images of the retrieved femurs with implants: (a) femurs stabilized with UFS intramedullary nail, (b) femurs stabilized with UFS-BG intramedullary nail. The axial cross-section images (1–4) correspond to the sites of hardtissue sections.

journal of the mechanical behavior of biomedical materials 40 (2014) 127 –139

135

UFS materials observed in the laboratory testing of the specimens.

3.3.

Mechanical testing of cadaver femurs with implants

The experiment with cadaver rabbit femurs stabilized with UFS implants showed that the creation of the standard-sized defect in the subtrochanteric region reduced the torsional strength of the bones to 34% of the torsional strength of the intact femur (Zhao et al., 2009). Insertion of the UFS intramedullary nails slightly improved the torsional strengths of the bones, to 45% of the torsional strength of the intact femur. However, the differences between the femurs with unsupported defects (Zhao et al., 2009) and the femurs stabilized with UFS intramedullary nails were not statistically significant.

3.4.

Animal study

All animals underwent uneventful functional recovery from surgery and were followed up for 12 weeks as planned. No infections were encountered and no adverse tissue reactions were observed at the retrieval of the bone specimens. However, in both implantation groups (UFS and UFS-BG), half of the femurs showed signs of undisplaced healed fractures at the site of the original defect. Results of the biomechanical testing of the retrieved femurs are shown in Tables 4 and 5. In the intra-animal comparison of the retrieved femurs (Table 4), there were statistically significant differences between the healed femurs with implants and the contralateral intact femurs. In both implantation groups (UFS and UFS-BG), the maximum torque values were lower for the healed femurs with implants compared with the contralateral femurs. In turn, the torsional stiffness and the angle of rotation were higher for the healed femurs with implants compared with the contralateral intact femurs (Table 4). However, no statistically significant differences in the results of mechanical testing of femurs stabilized with UFS and UFS-BG intramedullary nails were detected (Table 5). In both implantation groups, the average maximum torque of the healed femurs was around 83% of the value measured for the contralateral femurs. Likewise, for the healed femurs with implants, the average torsional stiffness and the angle of rotation were around 74% and 124% of the values measured for the contralateral femurs, respectively (Table 5). Micro-CT-based reconstructions of the femurs stabilized with FRC intramedullary nails revealed a good level of bone healing of the cortical bone defects (Fig. 4a). In both implantation groups (UFS and UFS-BG), the newly formed peri-implant bone was observed in micro-CT imaging (Fig. 5a and b). In histological examination, in both implantation groups (UFS and UFS-BG), new peri-implant bone formation was observed at all sites selected for the analysis (Fig. 6a, b). No implant-related debris formation or adverse tissue reactions were observed. In the group of femurs stabilized with UFS intramedullary nails, most of the newly formed peri-implant bones were seen in sections 1 and 4. Typically, there was an interlayer of fibrous tissue between the surface of the implant and the newly formed bone. The extreme case is shown in Fig. 7a. However, frequently, the fibrous interlayer was thin (Fig. 7b). Moreover, in some cases,

Fig. 6 – Histological sections at the selected anatomical locations: (a) femurs stabilized with UFS intramedullary nail, (b) femurs stabilized with UFS-BG intramedullary nail. Bone is stained in red color (van Gieson). New bone formation is seen around both types of the intramedullary implants. there was a direct contact between the newly formed bone and the surface of the implant (Fig. 7c). In the group of femurs stabilized with UFS-BG intramedullary nails, most of the new bone was observed in sections 1 and 4. At the locations of the surface-bound BG granules, no fibrous capsule was observed in any of the sections. As expected, BG granules acted as sites for bone attachment. However, there was a gradual decrease in the amount of the surface-bound BG granules from the most proximal section 1 to the most distal section 4 (Fig. 7d–f). Moreover, BG granules detached from the implant surface were found in abundance in section 1 (Fig. 7d). These granules could have detached from the implant surface during the implantation. In addition, there were isolated cases of BG granules not promoting new bone formation (Fig. 7f).

4.

Discussion

Based on the encouraging results of our previous study (Zhao et al., 2009), this study was performed as an attempt to

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Fig. 7 – Histological images of the implant–tissue interface. Arrows indicate implant surface, the marker (n) indicates surfacebound BG granules, the marker (nn) indicates detached BG granules. NB¼ newly-formed bone. (a) A case of implant encapsulation by fibrous tissue (UFS). (b) A case of implant incorporation with bone (UFS). (c) A close up view demonstrating direct peri-implant bone opposition to the implant surface (UFS). (d) Bone formation adjacent to the BG granules (UFS-BG); note the completely reacted granules with typical bone formation from inside the granule. (e) Newly formed bone directly attaching to surface-bound BG granules (UFS-BG). (f) BG granules with and without bone attachment (UFS-BG surface).

improve the torsional strength of the FRC intramedullary nails made of BisGMA/TEGDMA-reinforced with unidirectional E-glass fibers. The addition of a bidirectional fiber sleeving to the unidirectional fiber core of reinforcement component of the composite improved the torsional strength of the UFS specimens at the expense of the flexural strength. No expected improvement in the in vivo performance of the UFS implants could be confirmed. The occurrences of undisplaced fractures in the defect area of the rabbit femurs indicated that the mechanical properties of the UFS implants remained suboptimal in prevention of fractures. In this study, the mechanical properties of two implant materials, UF and UFS, were assessed in three-point bending and torsion tests. For the UF specimens, the bending stress at the start of surface cracking, σf1 ¼745.5 MPa, the flexural modulus of the UF specimens, EB ¼17 GPa, was close to that of human cortical bone (12.5 GPa) (Lotz et al., 1991). The shear stress at the start of surface cracking was around 25.5 MPa (Tables 2 and 3). Previously reported mechanical properties of intramedullary nails made of BisGMA/TEGDMA-reinforced with unidirectional E-glass fibers were inferior to the properties of the UF specimens (Zhao et al., 2009). The flexural strength was 430 MPa, the flexural modulus was 12 GPa and the torsional strength was around 14 MPa (Zhao et al., 2009). The differences between those materials and the UF assessed in this study are probably related to the use of a different type of E-glass fibers pre-impregnated with a more cross-linked resin. Nevertheless, in our previous study, the FRC intramedullary nails provided sufficient structural support for the healing of the bone defects in most of the animals, while

healed peri-implant fractures occurred in two animals (Zhao et al., 2009). FRC intramedullary nails made of liquid crystalline polymer reinforced with long unidirectional carbon-fiber (LCP/CF) were studied in rabbit implantation models (Kettunen et al., 1998a, 1998b; Kettunen et al., 1999). The reported flexural strength of LCP/CF implants was 448 MPa, the flexural modulus was 43 GPa (Kettunen et al., 1998a, 1998b) and the torsional strength was 19.4–24.4 MPa (Suokas et al., 1993). Limited mechanical integrity of LCP/CF intramedullary nails and a number of non-unions were observed in a metaphyseal osteotomy model in rabbit (Kettunen et al., 1999). The mechanical properties of the UF specimens made of BisGMA/TEGDMA-reinforced with unidirectional E-glass fibers were superior to those of LCP/CF. It should be noted, that regardless of the values of maximum flexural and shear stresses (flexural and torsional strengths), the values of flexural and shear stresses at the start of surface cracking were of primary importance in this study. As in reality, these values were expected to indicate the failure of the FRC implants. The addition of a bidirectional fiber sleeving to the reinforcement component of the composite signified marked differences in the mechanical properties of the UF and UFS specimens. While the shear stress at the start of surface cracking increased to 53.5 MPa; the flexural stress at the start of surface cracking dropped to 278.2 MPa (Tables 2 and 3). For the UFS specimens, the shear stress at the start of surface cracking was superior to that of the UF specimens. In addition, the torsional strength of LCP/CF was lower (Suokas et al., 1993). The flexural strength of UFS specimens was relatively high (602 MPa), still inferior to that

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of UF specimens (802 MPa) (Table 2). In addition, there were marked differences in the values of flexural stress at the start of surface cracking measured for UFS specimens (278 MPa) compared with that of UF specimens (745.5 MPa) (Table 2). The geometry of the FRC intramedullary nails is determined by the shape of the intramedullary cavity of the rabbit femur (Fig. 2a). Consequently, regardless of the configuration of the reinforcement component, the cross-sectional area of the implants is fixed. In turn, the addition of the bidirectional fiber sleeving in the UFS specimens led to the decrease in the amount of the unidirectional core fibers in the reinforcement component (Fig. 1a and b). Apparently, with the fixed crosssectional area, there is a tradeoff between the torsional and the flexural strength of the implants. Let us compare the results of the mechanical testing of the UFS specimens with the results of the biomechanical testing of the femurs stabilized with the FRC intramedullary nails (Tables 4 and 5). When tested in torsion, the femurs with implants failed at the rotational angle of around 171. The intact femurs failed at around 231. In turn, UFS specimens tested in torsion with the gauge length of 35 mm started to exhibit surface cracking at the rotational angle of around 291. It appears that the FRC intramedullary nails can rotate more than the bone. This finding suggests that implants may not have provided sufficient torsional support to the bones. One possible way to improve the mechanical properties of the intramedullary nails would be to completely or partially substitute the types of reinforcement fibers, e.g. S-glass fibers. However, this needs confirmation by mechanical tests and FE analysis. One possible implication could be that rabbit is not an ideal animal to test the composite implants intended for use in humans. Therefore, a larger animal model, such as dog or minipig, would be preferable in future studies. In this study, the FRC specimens and implants were handmade. As a result, the presence of voids was observed in micro-CT images and histological sections (Figs. 5, 6 and 7). The voids resulted from air bubbles left in the resin after the polymerization. These voids might have influenced the results of mechanical tests and the in vivo performance of the implants. The control of the quality of the specimens was performed by micro-CT imaging. However, low attenuation of X-rays by resin matrix resulted in difficulties in distinguishing between the voids and the excessive resin matrix. In general, the volume of zones of low attenuation inside the implants was less than 10%. Obviously application of a well-controlled fabrication method could help in preparation of more uniform specimens and implants. As an example, filament winding can provide tight winding of the bias fibers around the unidirectional core fibers. In turn, pressure molding could be applied for the removal of the voids. It was reported that in contrast to metallic implants, less stiff composite implants may produce substantial mechanical stresses at the bone-implant interfaces (Huiskes, 1993; Evans and Gregson, 1998). Therefore, bioactive fixation of load-bearing implants may be required to tackle with this issue (Evans and Gregson, 1998). BG granules have been clinically used as bone graft substitute (Lindfors et al., 2010; Frantzen et al., 2011; Heikkilä et al., 2011). In addition, BG granules were applied as a surface component of FRC implants tested in calvarial defect models of the rabbit

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(Tuusa et al., 2007; Tuusa et al., 2008). Therefore, surfacebound BG granules could provide such bioactive fixation in load-bearing applications if the quantity of BG is sufficient and the mechanical cohesion with resin matrix is strong. However, the use of the BG particulate as a surface component of FRC implants could be challenging (Zhao et al., 2009). In our previous study, BG granules in the fraction of 315– 500 mm were embedded in the surface of the FRC implants. Results indicated that the large number of BG granules was trapped inside the resin matrix and the surface area of the exposed granules remained limited (Zhao et al., 2009). Therefore, a smaller fraction of BG granules, 90–315 mm, was used in this study to increase the surface area of BG available for bone-bonding. To further increase the BG area available for bone attachment, the granules were not embedded in the resin but rather “glued” to the surface of the implants. Despite the theoretical advantages of this surface modification, in practice, the selected method of application of BG granules did not provide the desired effect. Results indicated that the BG granules detached from the implant upon implantation and possibly later during the follow-up period (Fig. 7d). Thus, the incorporation of BG granules within the surface layer of FRC remains a technological challenge, especially when the surface of the implant is subjected to high shear forces during the implantation. The possible advantages of the carbon FRC implants over metallic implants are well acknowledged in the literature (Evans and Gregson, 1998). The potential of less rigid composite plate fixation of long bones was demonstrated in animal models (Tayton et al., 1982; Woo et al., 1984; Gillett et al., 1985; Jockisch et al., 1992). Carbon FRC was also reported as a promising candidate material for intramedullary use in loadbearing applications (Kettunen et al., 1999). Attempts have been made to use fracture fixation plates made of carbon FRC at different anatomical locations in humans (Tayton et al., 1982; Ali et al., 1990; Pemberton et al., 1992, 1994). Reports of implant failure (Allcock and Ali, 1997), poor implant design (Adam et al., 2002) or accumulation of debris (Gillett et al., 1985; Kettunen et al., 1999; Jockisch et al., 1992; Minovic et al., 2001) were the reasons why carbon FRC implants never made clinical success. However, as a proof of concept, low-modulus composite hip stems (Akhavan et al., 2006) and high stiffness carbon fiber-reinforced bone plates (Rohner et al., 2005) were reported as valuable options to metallic implants. Glass-fiber reinforced FRCs present one of the newly available options of composite implant materials. Our previous study (Zhao et al., 2009) demonstrated the potential of these FRCs for use as loadbearing orthopedic implants. However, extensive research effort is required to prove their advantage over metallic implants. It should be noted that due to the large number of design variables, composites present a more complex system than the monolithic metals (Evans and Gregson, 1998). In addition, there is a synergetic combination of several factors, i.e. biological, environmental, mechanical, chemical, surface science etc., which contribute to the success or failure of a medical implant (Kohn and Ducheyene, 1992). In this study, the reinforcement component comprised of a combination of unidirectional E-glass fibers with the bidirectional fiber sleeving improved the torsional strength of the FRC implants. While the flexural properties of the implants

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were impaired, under physiological loading conditions, the FRC implants were able to support the femurs with defects without structural failure (delamination) and debris formation, even in the cases of peri-implant fractures. All these fractures were healed at the time of the retrieval of the bones with implants. Hence, the glass fiber reinforced composites assessed in this study could be promising load-bearing implant materials, provided the structural optimization of the reinforcement component is achieved.

5.

Conclusions

The reinforcement component comprised of a combination of unidirectional glass fibers with the bidirectional fiber sleeving improved the torsional strength of the FRC implants at the expense of the flexural strength. The FRC intramedullary nails were able to support the femurs with defects without structural failure (delamination) and debris formation, even in the cases of peri-implant fractures. Implants made of the glass fiber reinforced composites could be promising loadbearing implant materials, provided the structural optimization of the reinforcement component is achieved.

Acknowledgments This study was supported by the Finnish Funding Agency for Technology and Innovation (Tekes) (Grants nos. 40225/05, 40170/06) via Commercialization of Biomaterials Technology Program (COMBIO) and the EU Project for SMEs (6th EU Framework Program) NEWBONE (NMP3 CT-2007-026279 Contract no. 026279). Study was part of the activity of BioCity Turku Biomaterials Research Program (http://www.biomater ials.utu.fi).

references

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